Implantable medical devices have become increasingly more common over the last 50 years and have found applications in nearly every branch of medicine. Examples include joint replacements, vascular grafts, heart valves, ocular lenses, pacemakers, vascular stents, urethral stents, and many others. However, regardless of the application, implantable medical devices must be biocompatible, that is, they must be fabricated from materials that will not elicit an adverse biological response such as, but not limited to, inflammation, thrombogensis or necrosis. Thus, early medical devices were generally fabricated from inert materials such as precious metals and ceramics. More recently, stainless steel and other metal alloys have replaces precious metals and polymers are being substituted for ceramics.
Generally, implantable medical devices are intended to serve long term therapeutic applications and are not removed once implanted. In some cases it may be desirable to use implantable medical devices for short term therapies. However, their removal may require highly invasive surgical procedures that place the patient at risk for life threatening complications. Therefore, it would be desirable to have medical devices designed for short term applications that degrade via normal metabolic pathways and are reabsorbed into the surrounding tissues.
One of the first bioresorbable medical devices develop was the synthetic absorbable suture marketed as Dexon in the 1960s by Davis and Geck, Inc. (Danbury, Conn.). Since that time, diverse biodegradable polymer-based products have found acceptance as implantable medical devices and implantable medical device coatings, thereby alleviating the need for secondary invasive procedure(s) to remove implanted medical device(s).
Biodegradable polymers can be either natural or synthetic. In general, synthetic polymers offer greater advantages than natural materials in that they can be tailored to give a wider range of properties and more predictable lot-to-lot uniformity than can materials from natural sources. Synthetic polymers also represent a more reliable source of raw materials, ones free from concerns of immunogenicity.
In general, polymer selection criteria for use as biomaterials is to match the mechanical properties of the polymer(s) and degradation time to the needs of the specific in vivo application. The factors affecting the mechanical performance of biodegradable polymers are those that are well known to the polymer scientist, and include monomer selection, initiator selection, process conditions and the presence of additives. These factors in turn influence the polymer's hydrophilicity, crystallinity, melt and glass-transition temperatures, molecular weight, molecular-weight distribution, end groups, sequence distribution (random versus blocky) and presence of residual monomer or additives. In addition, the polymer scientist working with biodegradable materials must evaluate each of these variables for its effect on biodegradation. Known biodegradable polymers include, among others, polyglycolide (PGA), polylactide (PLA) and poly(ε-caprolactone) (PCA). However, these polymers are generally hydrophobic and their structures are difficult to modify. Consequently, the polymer's physical characteristics are difficult to modify, or tune, to match specific clinical demands. For example, polymers made from PLA are extremely slow to degrade and thus not suited for all applications. To address this deficiency polymer scientists have developed co-polymers of PLA and PCA. However, biodegradation rates remain significantly limited.
Additionally, recent advances in in situ drug delivery has led to the development of implantable medical devices specifically designed to provide therapeutic compositions to remote anatomical locations. Perhaps one of the most exciting areas of in situ drug delivery is in the field of intervention cardiology. Vascular occlusions leading to ischemic heart disease are frequently treated using percutaneous transluminal coronary angioplasty (PTCA) whereby a dilation catheter is inserted through a femoral artery incision and directed to the site of the vascular occlusion. The catheter is dilated and the expanding catheter tip (the balloon) opens the occluded artery restoring vascular patency. Generally, a vascular stent is deployed at the treatment site to minimize vascular recoil and restenosis. However, in some cases stent deployment leads to damage to the intimal lining of the artery which may result in vascular smooth muscle cell hyperproliferation and restenosis. When restenosis occurs it is necessary to either re-dilate the artery at the treatment site, or, if that is not possible, a surgical coronary artery bypass procedure must be performed.
Recently, it has been determined drug-eluting stents coated with anti-proliferative drugs such as, but not limited to, rapamycin and its analogs and paclitaxel have shown great promises in preventing restenosis. However, there is a need to develop additional and potentially more efficacious drug-eluting stents (DES). One critical factor in DES efficacy is the drug elution rate. Drug-elution is generally a factor of the drug's solubility in the polymer coating applied to the stent.
Presently, bio-stable, that is non-resorbable polymers, are used as drug eluting coatings for metal stents. The polymer scientist has many different options when selecting a suitable bio-stable polymer and recently several of the present inventors have made significant advances in turning polymer coatings useful for drug elution (see co-pending U.S. patent application Ser. No. 11/005,463). However, the number and type of bioresorbable polymers is much more limited. Therefore, in order for new bioresorbable polymeric medical devices to be developed which have the same functional diversity as their biostable polymer counterparts, it is necessary to first develop new and useful bioresorbable polymers that can be tuned to match drug solubility and provide greater control over resorbability rates.